Biodegradable extended release microsphere-hydrogel ocular drug delivery system and method

ABSTRACT

A hydrogel delivery composition and method, including degradable microcapsules suspended in a degradable thermo-responsive hydrogel. The hydrogel is thermo-responsive at a physiological temperature and changes after application to a more solid state due to body temperatures. The composition includes one or more treatment agents to be released over time as the composition degrades. The composition can be varied to modify the structure and/or release of the treatment agent. The degradable microcapsules include one or more of magnesium hydroxide (Mg(OH)2), bovine serum albumin (BSA), polyethylene glycol (PEG), and sucrose to improve release duration.

CROSS REFERENCE TO RELATED APPLICATION

This application is a continuation-in-part application of Ser. No.16/845,184, filed on 10 Apr. 2020, which claims the benefit of U.S.Provisional Patent Application, Ser. No. 62/832,977, filed on 12 Apr.2019, and is a continuation-in-part application of Ser. No. 15/273,098,filed on 22 Sep. 2016, which claims the benefit of U.S. ProvisionalPatent Application, Ser. No. 62/232,545, filed on 25 Sep. 2015. Theco-pending parent application(s) and provisional patent application(s)are hereby incorporated by reference herein in their entirety and ismade a part hereof, including but not limited to those portions whichspecifically appear hereinafter.

FIELD OF THE INVENTION

This invention relates generally to drug or other treatment deliveryand, more particularly, to a hydrogel delivery composition and methodthat can deliver treatment materials in a controlled matter for anextend period of time.

BACKGROUND OF THE INVENTION

Choroidal neovascularization (CNV) secondary to age-related maculardegeneration (AMD) is a leading cause of vision loss in elderly patientsin developed nations. It has been demonstrated that vascular endothelialgrowth factor (VEGF) plays a key role in the pathophysiology of thedisease. The U.S. Food and Drug Administration has approved severalanti-VEGF therapeutics to treat CNV secondary to AMD, namely ranibizumab(Lucentis®, Genentech, South San Francisco, Calif., USA), aflibercept(Eylea®, Regeneron, Tarrytown, N.Y., USA) and brolucizumab (Beovu®,Novartis, Basel, Switzerland). Additionally, the use of anti-VEGFtherapeutics has increased for other indications such as diabeticmacular edema (DME) and vein occlusion. Due to these successfuloutcomes, it is not surprising that there has been a paradigm shift inthe use of pharmacological therapeutics over surgical techniques duringthe last decade to treat visually devastating diseases.

The current standard to deliver anti-VEGF therapeutics is through amonthly (or bimonthly) bolus intravitreal injection. With eachinjection, serious potential complications can arise includingendophthalmitis, retinal detachment, intravitreal hemorrhage, andcataract. To improve upon the socio-economic impact associated withrepeated injections and to further lower the risk of associatedpotential complications, there is a need to reduce the number andfrequency of injections.

A multitude of drug delivery systems (DDSs) have been successful incontrolling and extending the release of various model drugs. A commontheme among these DDSs is the use of polymeric microspheres to controlrelease. However, when microspheres are injected into the eye asindependent units, they can become lodged in ocular tissues (e.g.,trabecular meshwork), which may cause unintended and seriouscomplications. Thus, there is a continuing need for improved drugdelivery systems, particularly in the development of intraocular DDSs.

SUMMARY OF THE INVENTION

A general object of the invention is to provide an improved deliverysystem and method for delivering treatment material(s), such as on orwithin the eye.

The general object of the invention can be attained, at least in part,through a delivery composition including degradable microcapsules (e.g.,nano- or micro-spheres or other similar microencapsulation structures)suspended in a degradable thermo-responsive hydrogel. The hydrogel isthermo-responsive in that it desirably changes its physical state from aliquid-like state at room temperature to a more solid state at bodytemperature (e.g., at a physiological temperature of about 32° C. toabout 37° C.). The microcapsules encapsulate and release a treatmentagent.

The rate of release can be controlled via the rate of degradation, whichcan be controlled, for example, by the components of the microcapsules,altering a chemical component ratio, varying cross-linking, themolecular weight of the polymer(s) used, surface modification, and/ormanufacturing procedures, such as the organic solvent used. In someembodiments a non-encapsulated treatment agent can optionally beadditionally dispersed within the hydrogel, and/or microcapsules havingat least two different release rates can be used together.

Embodiments of this invention focus on anti-VEGF as a treatment agent,but the invention can also be used to deliver other drugs, such asophthalmic drugs (for example, antibiotics for ocular surgery),anti-platelet-derived growth factor (anti-PDGF) agents, cells, deliverycells, a corticosteroid, enzymes, peptides, nucleic acids, orcombinations thereof. The invention can, for example, reduce oreliminate patient administration of treatment after ocular surgery, suchas eliminating the need to administer topical (eyedrop) antibiotics forseveral days. Most often, the patient compliance is low and at risk forinfection.

Embodiments of this invention include a delivery composition with atreatment agent microencapsulated in degradable microcapsules that aresuspended in a degradable thermo-responsive hydrogel. Biodegradablemicrocapsules can be, for example, based on poly(lactic-co-glycolicacid) (PLGA). The degradable microcapsules desirably further include orare formed with magnesium hydroxide (Mg(OH)₂) and bovine serum albumin(BSA), to provide extended duration, as long as 6 months or more. Themicrocapsules can include 0.001% to 20% w/v BSA and 0.001% to 9% w/vMg(OH)₂.

Microcapsules of embodiments of this invention can alternatively oradditionally include polyethylene glycol (PEG) and sucrose. Currentlypreferred embodiments in 0.001% to 20% w/v PEG and 0.001% to 10% w/vsucrose, alone or preferably with 0.001% to 20% w/v BS and/or 0.001% to9% w/v Mg(OH)₂.

The invention further includes methods of forming microencapsulatedtreatment agents. The methods include varying the microcapsulecomponents, amounts, and/or crosslinking, and adjusting the formationprocess. For example, the type of organic solvent used can impactdegradation duration.

The invention further includes a delivery composition having a treatmentagent microencapsulated in degradable microcapsules suspended in adegradable thermo-responsive hydrogel. The degradable microcapsules areformed of or otherwise include in the micro-structure at least two of:polyethylene glycol (PEG), magnesium hydroxide (Mg(OH)₂), bovine serumalbumin (BSA), and sucrose. The hydrogel is thermo-responsive at aphysiological temperature of about 32° C. to about 37° C. to provide aliquid-like state at room temperature and more solid state at bodytemperature.

Embodiments of this invention further include a drug delivery systemwhere two or more different microcapsules (e.g., two or more differentreleases and/or two or more different treatment agents) made accordingto this invention are suspended in a hydrogel. The two differentmicrocapsules can be two different nanoparticles, two differentmicroparticles, or combinations of nanoparticles and microparticles. Inembodiments of this invention, nanoparticles or microparticles with afirst treatment agent and nanoparticles or microparticles with adifferent second treatment agent are suspended in a same hydrogel.Additional treatment agents can be again also included in the hydrogel.A mix of nanoparticles and microparticles in hydrogel are beneficial,for example, for combinations of hydrophobic drugs and biologics. Smallmolecules or hydrophobic drugs (e.g., dexamethasone) can be used innanoparticles, as well as other agents like antibiotics or othersteroids. Microparticles are better suited in some embodiments forprotein-based drugs and delicate biologics (e.g., AFL, GDF5, ENT), butother agents or cells can be encapsulated in microparticles.

In embodiments of this invention, the hydrogel includes a suspendedfirst plurality of nanoparticles or microparticles having a first size,a first release time, and/or a first treatment agent selected from ananti-VEGF agent, an anti-PDGF agent, cells, delivery cells, anantibiotic, a corticosteroid, enzymes, peptides, nucleic acids, orcombinations thereof (e.g., any specific agent discussed herein). Thehydrogel further includes a suspended second plurality of nanoparticlesor microparticles having a different second size, second release time,and/or second treatment agent independently selected from an anti-VEGFagent, an anti-PDGF agent, cells, delivery cells, an antibiotic, acorticosteroid, enzymes, peptides, nucleic acids, or combinationsthereof (e.g., any specific agent discussed herein).

In embodiments of this invention, exemplary treatment agents includenonsteroidal anti-inflammatory agents (e.g., bromfenac, nepafenac),antibiotics (e.g., vancomycin, fluoroquinolones (like moxifloxacin),glaucoma treatments (beta blockers, alpha agonist, prostaglandin analogand carbonic anhydrase inhibitors), and combinations thereof. Exemplarycombination treatments for use according to this invention include,without limitation, carbonic anhydrase inhibitors and beta blocker forglaucoma (e.g., dorzolamide and timolol), alpha agonist and beta blockerfor glaucoma (e.g., brimonidines and timolol), carbonic anhydraseinhibitor and alpha agonist (e.g., brinzolamide and brimonidine),moxifloxacin and dexamethasone for infection and inflammation, and/orneomycin and polymyxin and dexamethasone (a 3 drug combination) forinfection and inflammation.

The invention further comprehends a method of delivering a compound toan eye, including: applying to or into an eye of a mammal a compositionin a first physicochemical state, wherein the composition comprises amicroencapsulated treatment agent suspended in a thermo-responsivehydrogel; and the composition changing to a second physicochemical stateupon administration, wherein the second physicochemical state is moresolid than the first physicochemical state and degradable to release themicroencapsulated treatment agent. The applied microcapsules release thebioactive treatment agent over time after application.

Embodiments of this invention include a method of delivering a compoundto an eye, including steps of: applying to or into an eye of a mammal acomposition in a first physicochemical state, wherein the compositioncomprises a treatment agent microencapsulated in degradablemicrocapsules suspended in a degradable thermo-responsive hydrogel,wherein the degradable microcapsules are formed including at least twoof: polyethylene glycol (PEG), Mg(OH)₂, bovine serum albumin, andsucrose; and the composition changing to a second physicochemical stateupon administration, wherein the second physicochemical state is moresolid than the first physicochemical state, wherein the degradablemicrocapsules release the microencapsulated treatment agent over timeafter applying.

The method can be used with any composition, and microcapsulecomponents/amounts, described herein. In embodiments of this invention,the degradation can be controlled by the type of and/or amount ofcrosslinking within the composition, thereby controlling the release themicroencapsulated treatment agent. Using more than one microencapsulatedtreatment agent, such as different microencapsulations for a same agenttype, and/or also placing non-encapsulated agent in the hydrogel itself,can also or further be used to provide a desired release.

Compositions of this invention can be applied by any suitable manner,desirably in a first physicochemical state, such as by intravitrealinjection, by periocular or transcleral injection, by topicalapplication including scleral structures, by intracameral application,by suprachoroidal application, within ocular implants, or combinationsthereof. Injections can be made by small gauge needle (e.g., 20 gauge to33 gauge) or by microcatheter.

Other objects and advantages will be apparent to those skilled in theart from the following detailed description and examples taken inconjunction with the appended claims and drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows: A) a representative SEM image of ranibizumab-loadedmicrospheres (scale bar: 2 μm); B) a ranibizumab-loaded DDS in a 28Gneedle; C) a representative SEM image of aflibercept-loaded microspheres(scale bar: 2 μm); and D) an aflibercept-loaded DDS in a 28G needle atroom temperature.

FIG. 2 includes graphs of: A) cumulative percent-release of ranibizumabfrom microspheres (“alone”) and microspheres suspended in hydrogel (“inhydrogel”), where the microspheres in hydrogel had a significantly lowerD3 and a steadier, extended release; and B) cumulative drug-release frommicrospheres in hydrogel, where after Day 7, 0.153 μg/day of ranibizumabwas released and an additional, 25.1±12.1% (27.3 μg) of encapsulateddrug remained entrapped in the hydrogel after complete degradation ofthe microspheres (circled). Error bars represent standard error (n=3).

FIG. 3 includes graphs of: A) cumulative percent-release of afliberceptfrom microspheres (“alone”) and the effect of suspending microspheres inhydrogel (“in hydrogel”), where aflibercept-loaded microspheressuspended in hydrogel had a significantly lower D3 and a steadier,extended release; and B) cumulative drug-release from microspheres inhydrogel, where after the first week, 0.065 μg/day of aflibercept wasreleased, and an additional, 32.7±1.1% (21.4 μg) of encapsulated drugremained entrapped in the hydrogel after complete degradation of themicrospheres (circled in red). Error bars represent standard error(n=3).

FIG. 4 shows bioactivity of the ranibizumab-loaded DDS (“Rani.-gel”).Significant bioactivity (p<0.05) is seen throughout release ofranibizumab compared to the drug-free DDS (“Blank-gel”; right). Errorbars represent standard error (n=16, left; n=6, right).

FIG. 5 shows bioactivity of the aflibercept-loaded DDS (“Aflib.-gel”).Significant bioactivity (p<0.05) is seen throughout release ofaflibercept compared to the drug-free DDS (“Blank-gel”; right). Errorbars represent standard error (n=16, left; n=6, right).

FIG. 6 shows release of ranibizumab from biodegradable microspheres andbiodegradable thermo-responsive hydrogel, where different amounts ofcrosslinker were used to change the degradation rate.

FIG. 7 shows an example of a CNV lesion area quantification technique:A) the original image opened in ImageJ and three levels are selected.After applying MOT, three images are produced: B) Region 0 (background),C) Region 1 (diffuse leakage), and D) Region 2 (CNV). With Region 2selected, the threshold is adjusted to include all pixels within theimage. The lesion is outlined (square in E) and the number of pixelswithin the lesion is measured, which is then converted into an area.

FIG. 8 shows average single flash ERG intensity-response functions fornon-treated animals (top, a-wave; bottom, b-wave). Lines are the resultsof the Naka-Rushton analysis to the cumulative pool of data points andsymbols represent different time points of investigated time frame. Nosignificant differences were seen in either maximal a- and b-waveresponse or half-maximal a- and b-wave response.

FIG. 9 shows average single flash ERG intensity-response functions fordrug-free DDS treated animals (top, a-wave; bottom, b-wave). Lines arethe results of the Naka-Rushton analysis to the cumulative pool of datapoints and symbols represent different time points of investigated timeframe. A significant difference in maximal b-wave response is seenbetween Control and Week 12 (p<0.05). No significant difference inhalf-maximal a- or b-wave response was observed.

FIG. 10 shows average single flash ERG intensity-response functions forranibizumab-loaded DDS treated animals (top, a-wave; bottom, b-wave).Lines are the results of the Naka-Rushton analysis to the cumulativepool of data points and symbols represent different time points ofinvestigated time frame. No significant differences were seen in eithermaximal a- and b-wave response or half-maximal a- and b-wave response.

FIG. 11 shows average single flash ERG intensity-response functions forbolus ranibizumab treated animals (top, a-wave; bottom, b-wave). Linesare the results of the Naka-Rushton analysis to the cumulative pool ofdata points and symbols represent different time points of investigatedtime frame. No significant differences were seen in either maximal a-and b-wave response or half-maximal a- and b-wave response.

FIG. 12 shows average single flash ERG intensity-response functions foraflibercept-loaded DDS treated animals (top, a-wave; bottom, b-wave).Lines are the results of the Naka-Rushton analysis to the cumulativepool of data points and symbols represent different time points ofinvestigated time frame. A significant difference in maximal a-waveresponse is seen between Control and Week 1-Week 12 (p<0.05). Nosignificant difference in half-maximal a- or b-wave response wasobserved.

FIG. 13 shows average single flash ERG intensity-response functions forbolus aflibercept treated animals (top, a-wave; bottom, b-wave). Linesare the results of the Naka-Rushton analysis to the cumulative pool ofdata points and symbols represent different time points of investigatedtime frame. A significant difference in maximal a-wave response is seenbetween Control and Week 1-Week 12 (p<0.05). Additionally, a significantdifference in maximal b-wave response is seen between Control and Week1-Week 12 (p<0.05). No significant difference in half-maximal a- orb-wave response was observed.

FIG. 14 shows CNV lesion areas for all treatment groups. Although CNV isassumed to not be fully developed at Week 1, it is presented forcompleteness: A) No significant difference is seen between non-treatedanimals and drug-free DDS treated animals (“Blank DDS”) at any timepoint; B) ranibizumab-loaded DDS treated animals exhibited a significantdecrease in CNV lesion areas from Week 4-Week 12 compared to non-treatedanimals whereas bolus ranibizumab treated animals only had a significantdecrease in lesion area at Week 12; and C) bolus aflibercept treatedanimals had significantly smaller CNV lesions compared to non-treatedanimals at Week 4 and Week 12. However, aflibercept-loaded DDS treatedanimals had significantly smaller CNV lesions than bolus treated animalsfrom Week 8-Week 12 (p<0.05).

FIG. 15 shows representative composite FA images from Week 12 oftreatment groups: A) non-treatment, B) drug-free DDS, C) ranibizumabbolus, D) ranibizumab-loaded DDS, E) aflibercept bolus, and F)aflibercept-loaded DDS. As is seen, 5-6 CNV lesions were induced in eachtreatment group and the severity of the lesions were approximately thesame at Week 12.

FIG. 16 is a chart summarizing aflibercept recover percentage, accordingto a provided example.

FIG. 17 is a chart summarizing aflibercept stability, according to aprovided example.

FIG. 18 shows in vitro release of dexamethasone fromnanoparticle-hydrogel DDS with 20, 40, 60, and 80 mg/ml loading doses.Combination releases for 40, 60, and 80 mg/ml loading dose only have asingle replication.

FIG. 19 summarizes cumulative releases of aflibercept from single andcombination DDS. Error bars represent standard error (n=3).

FIG. 20A illustrates cumulative releases of DEX from DEX-singlemonotherapy DDS and combo-DDS, according to an example of thisinvention.

FIG. 20B illustrates cumulative release of AFL from AFL-singlemonotherapy DDS and combo-DDS, according to an example of thisinvention.

FIG. 21A shows an ETN interval release profile for an ETN (10 mg/ml) DDSand ETN (10 mg/ml)+GDFS (20 mg/ml) DDS, according to an example of thisinvention.

FIG. 21B shows an ETN cumulative release profile for an ETN (10 mg/ml)DDS and ETN (10 mg/ml)+GDFS (20 mg/ml) DDS, according to an example ofthis invention.

FIG. 21C shows an GDFS interval release profile for an GDFS (30 mg/ml)DDS and ETN (10 mg/ml)+GDFS (20 mg/ml) DDS, according to an example ofthis invention.

FIG. 21D shows an GDFS cumulative release profile for an GDFS (30 mg/ml)DDS and ETN (10 mg/ml)+GDFS (20 mg/ml) DDS, according to an example ofthis invention.

FIG. 22 shows an average VAN release from a PNIPAAm-PEG-DA (MW 575)hydrogel, shown as average percent cumulative VAN release from a 1-mlthermo-responsive hydrogel on the primary Y-axis and VAN release termsof the average cumulative amount of drug released at each point on thesecondary Y-axis, according to an example of this invention

DESCRIPTION OF THE INVENTION

The present invention provides a hydrogel delivery composition andmethod that can delivery treatment materials (e.g., drugs, etc.) in acontrolled matter for an extend period of time. Encapsulating andreleasing, for example, a bioactive drug for a long period of time is adifficult task. Many laboratories have tried to do this with limitedsuccess.

The composition of embodiments of this invention combines both hydrogeland microcapsules for treatment. Degradable microcapsules are suspendedin a degradable thermo-responsive hydrogel. The hydrogel isthermo-responsive at a physiological temperature of about 32° C. toabout 37° C., such as to change, via crosslinking, from a liquid-likestate at room temperature to a more solid state at body temperature.Exemplary hydrogel materials include poly(N-isopropylacrylamide),poly(lactic acid), polysaccharide chitin, alginate, polyethylene glycol(PEG) and/or diacrylate (DA), or combinations or copolymers thereof.Exemplary copolymers include block copolymers PEG-PLLA-DA orPLA-PEG-PLA.

The thermo-responsive hydrogels of this invention provide an injectableplatform material that solidifies once injected into position. Smallgauge needles, e.g., about 20 to 33 gauge, and preferably about 25 to 27gauge, such as used in conventional intravitreal injection treatment,can be used to deliver the composition of this invention. This can be asignificant benefit as the injections can occur in doctor's officerather than in an operating room. The thermo-responsive hydrogels ofthis invention are injected in the “liquid-like” form and “solidify” toa gel material at body temperature. The transition time can becontrolled, so that the hydrogel can be used for various applications.

The hydrogels are desirably biodegradable, and can be either partiallyor fully biodegradable. The duration of degradation can also becontrolled. In embodiments of this invention, the degradation rate iscontrolled by the type and/or amount of crosslinker contained in the‘liquid’ form of the composition. In another embodiment, the degradationrate is modified by using different concentrations of and/or chaintransfer agents (e.g., glutathione) with poly(N-isopropylacrylamide).

In embodiments of this invention, the thermo-responsive hydrogel is usedto suspend or otherwise entrap microcapsules, e.g., nano- ormicrospheres, for easier delivery as well as localized delivery of themicrocapsules. Microcapsules generally will not stay in the eye wheninjected alone. Studies have shown that microcapsules will clear withinabout fifty days in normal eyes and within fourteen days invitrectomized eyes. Injecting microcapsules alone to deliver a long-termdrug treatment will not work. Another concern is that “free floating”microcapsules (if they are big enough) can interfere with visualfunction, in that a patient may ‘see’ microcapsules floating andinterfere with vision. By incorporating microcapsules into hydrogelaccording to this invention, the hydrogel structure ‘keeps’ themicrocapsules in place to provide longer-term delivery. Keepingmicrocapsules in the hydrogel also avoids interfering with the centralvisual pathway (visual perception).

The microcapsules are also desirably biodegradable, and can be eitherpartially or fully biodegradable. The duration of degradation and/orrelease can also be controlled. The degradation time and thus releaserate can be controlled, for example, through changing the polymer ratio,modifying the molecular weight of polymer(s), surface modification,and/or manufacture procedure. Currently preferred biodegradablemicrocapsules are based on poly(lactic-co-glycolic acid) (PLGA).Exemplary microcapsules according to this invention are formed using anysuitable microencapsulation process from poly(lactic-co-glycolic acid),poly(lactic acid) (PLA), polysaccharide chitin, alginate(polysaccharide), or combinations or copolymers thereof. Exemplarycopolymers include, without limitation, PLGA-chitin and block copolymerswith polyethylene glycol (PEG) and/or diacrylate (DA), such as PLGA-PEG,PEG-PLLA-DA, PEG-PLGA-PEG, and PLGA-PEG-PLGA.

Microcapsules according to preferred embodiments of this inventioninclude one or more, and desirably at least two or more, enhancementadditives, in addition to the above component(s), selected frommagnesium hydroxide (Mg(OH)₂), bovine serum albumin (BSA), polyethyleneglycol (PEG) and sucrose. Without wishing to be bound by theory, BSA andMg(OH)₂ appear have an impact on large protein molecule drugs likeanti-VEGF for the extended release duration, whereas PEG and sucroseappear more important in an initial release phase.

Microparticles of embodiments of the present invention includecomponents of the following ranges: 0.001% to 20% w/v BSA, 0.001% to 9%w/v Mg(OH)₂, 0.001% to 20% w/v PEG, and/or 0.001% to 10% w/v sucrose. Inan exemplary embodiment, microcapsules include components of thefollowing amounts: about 10-14%, or about 12%, w/v %, BSA; about 2-4%,or about 3%, w/v, Mg(OH)₂; about 8-12%, or about 10%, w/v PEG(PEG-8000); and/or about 1.5-3.5%, or about 2.5%, w/v sucrose. As willbe appreciated the actual amounts per individual microcapsule in acollection can vary due to the formation process.

The composition includes a treatment agent microencapsulated indegradable microcapsules suspended in a degradable thermo-responsivehydrogel, for the purpose of delivering the treatment agent to adelivery site. The microencapsulated treatment agent-in-hydrogelcomposition of this invention changes to a second, more solidphysicochemical state upon administration, and then degrades to releasethe treatment agent. The hydrogel keeps the microcapsule in place(localized to delivery site) and prevents free movement of microcapsulein and/or around the eye. The hydrogel can also act as another releasebarrier and prolong the release time. The applied microcapsules releasethe bioactive treatment agent over time after application.

As discussed above, the hydrogels can be either fully biodegradable ornot. Biodegradable microspheres are generally based onpoly(lactic-co-glycolic acid) (PLGA). The duration of degradation canalso be controlled, such as, for example, by changing the polymer ratio,varying the cross-linking and/or the amounts/ratios of the enhancementadditives discussed above. During the fabrication steps, variousexcipients can be added to protect the protein-based drug. Thedegradation time can also be controlled by the organic solvent used formicrocapsule formation. The type and/or evaporation time of the organicsolvent can influence encapsulation and morphology of microparticles(which will influence drug release profiles and release duration). Therange of organic solvent evaporation time can be between 5 min to 6hours. Using preferred solvents such as dichloromethane (DCM) or ethylacetate (EA) as organic solvent has been found to enhance encapsulationand protein (drug) stability. Triacetin can also be used as an organicsolvent, but appears not as optimal as DCM or EA.

Any suitable treatment agent can be encapsulated by any suitableencapsulation method. Exemplary treatment agents include, withoutlimitation, drugs, antibiotics (e.g., vancomycin, gentamicin),corticosteroids (e.g., dexamethasone, triamcinolone acetonide),anti-VEGF (e.g., pegaptanib, bevacizumab, ranibizumab, aflibercept,brolucizumab, conbercept), anti-PDGF, peptides, enzymes, biologicalagents (e.g., nucleic acids: DNA, RNA), various stem or other cells, orcombinations thereof.

Treatment agent(s) can release from the microcapsule only, from both themicrocapsule and the hydrogel, from combinations of different sizemicrocapsules, from both different size of microcapsules and hydrogel,etc. Combination of treatment agents can be released (e.g., anti-VEGFand anti-PDGF, or combination of two different drugs) from the same ordifferent microcapsules and/or from the microcapsule and the hydrogel.These variations allow for different release rates and/or treatments. Inone embodiment of this invention, the thermo-responsive hydrogel is usedto deliver a same or different agent in addition to the microspheres.The hydrogel itself can be, for example, drug-loaded and the release canbe controlled; hence, providing another option to deliver agents inconjunction with microcapsules. This can be ideal in a long-termdelivery. For example, hydrogels can deliver “loading” doses of drugsand the microspheres deliver “maintenance” doses.

Microcapsule-hydrogel platforms of this invention can provide long-termrelease of anti-VEGF, such as a controlled release of anti-VEGF for over6 months. Both in vitro and in vivo studies have shown that releasedanti-VEGF agents are active for the duration. This means that anti-VEGFcan be safely stored in the drug delivery platform of this invention andthe released drug has a positive effect clinically. Both modifiedmicrocapsule fabrication steps (adding protective agent for treatmentagent) and using hydrogel provided both extended release and bioactivityof the drug.

The delivery system can encapsulate various ocular drugs. The platformof this invention can replace monthly conventional intravitrealinjection treatment. Another novel application is to release multipledrugs at the same time. By encapsulating different drugs (or the samedrug with different dosages), drugs can be controlled and released atdifferent rates. As an example, clinically, there is growing evidencethat dual treatment of anti-VEGF and anti-PDGF (both deliveredconventionally intravitreal injection method) may be beneficial overmonotherapy of anti-VEGF. The platform of this invention can be utilizedto release both anti-VEGF and anti-PDGF for an extended period of time.

As discussed above the composition of this invention can be injected,and can also be used topically or in other medical procedures/devices.As an example, an illuminated microcatheter (such as iTrackmicrocatheter, iScience Interventional) can be used to deliver variousagents (encapsulated drugs or cells) into the suprachoridal space.Microcatheter injection has not typically been successful due to abackflush of drug after the injection. The biodegradablethermos-responsive composition of this invention will help keep thetreatment agent in place (due to the change from liquid-like tosolid-gel) after injection.

The composition of this invention can also be injected around the eyeglobe to act as a temporary scleral structure, such as a scleral bucklefor retinal detachment treatment. Currently, only permanent scleralbuckles (made out of surgical sponge or plastic) are used even thoughthe original intent of buckle use was a temporary procedure. Thecomposition can be injected directly into the globe to providestructural support (like a scleral buckle). Because the composition isbiodegradable, there will be no need to remove the hydrogel.Furthermore, it can be combined with agents (e.g., antibiotic) toprevent any post-procedure infection. Since it will be simple injection,the procedure should be minimally invasive. Because the polymercomposition can be controlled, the hydrogel can be made more firm togive sufficient structural support.

The composition can be used in existing or future medical devices. Forexample, the composition can be filled inside of ocular implants such asa Port Delivery System (PDS, developed by ForSight Vision4 and licensedby Genentech), which is a refillable, non-biodegradable implant(surgical procedure), or other implant reservoirs. Patients typicallyneed 4-5 refills over 12 months. Because the composition of thisinvention can release for over 6 months, a combination system can extendtreatment.

The present invention is described in further detail in connection withthe following examples which illustrate or simulate various aspectsinvolved in the practice of the invention. It is to be understood thatall changes that come within the spirit of the invention are desired tobe protected and thus the invention is not to be construed as limited bythese examples.

EXAMPLES Example 1

A drug delivery system (DDS) was prepared, composed of biodegradablepoly(lactic-co-glycolic acid) (PLGA) microspheres suspended within athermo-responsive, injectable poly(N-isopropylacrylamide)(PNIPAAm)-based hydrogel. In the initial development of this DDS,ovalbumin was used as a model protein to demonstrate controlled andextended release for approximately 200 days. By suspending themicrospheres within hydrogel, the initial burst (IB) was greatly reducedand release was extended by ˜30% compared to microspheres alone. Severalexcipients were incorporated into the DDS to protect against the acidicdegradation products of PLGA as well as to protect the protein (drug)during preparation, storage, and release. The objective of this work wasto demonstrate the capability of the microsphere-hydrogel DDS of thisinvention to release either ranibizumab or aflibercept in a controlledand extended manner and to determine whether the drugs remain bioactivethroughout release.

Ranibizumab and aflibercept were radiolabeled (to determine releaserate) with iodine-125 using iodination beads (Pierce, Rockford, Ill.,USA) and then dialyzed against double-deionized water (ddH₂O) using adialysis cassette (MWCO 2 kDa, Pierce) to remove unbound, free iodine.Labeled proteins were lyophilized, weighed, and dissolved in 1×phosphate-buffered saline (PBS, pH 7.4) to create stock solutions of 10mg/mL for each protein, which were stored at −80° C.

All subsequent chemicals were purchased from Sigma-Aldrich (St. Louis,Mo., USA). Each anti-VEGF was incorporated into PLGA microspheres usinga modified double-emulsion, solvent evaporation technique described inOsswald C. R. et al., Controlled and Extended Release of a Model Proteinfrom a Microsphere-Hydrogel Drug Delivery System, Ann Biomed Eng., Apr.3, 2015. Briefly, each protocol contained the same excipients in theinner aqueous (w₁) and oil (o) phases and each protocol was made intriplicate. In the w₁ phase, 12.5 mg bovine serum albumin (BSA), 10 mgPEG (molecular weight, MW, 8 kDa) (added to ‘protect’ the agent), and2.5 mg sucrose were dissolved in 100 μL of stock anti-VEGF solution. Inthe o phase, 125 mg PLGA 75:25 (4-15 kDa, acid terminated) and 3.75 mgMg(OH)₂ (also added to ‘protect’ the agent) were dissolved in 0.5 mLdichloromethane.

The primary water-in-oil emulsion (w₁/o) was created by vortex at 3200rpm for 90 s (Fisher Scientific Analog Vortex Mixer; 120 V; speed 10×).The w₁/o do emulsion was immediately added to the outer aqueous phase(w₂) containing 10% (w/v) polyvinyl alcohol (PVA). The secondary(water-in-oil)-in-water emulsion (w₁/o/w₂) was created by vortex at 2200rpm (speed 5×) for 90 s. The w₁/o/w₂ emulsion was then added to 75 mL of0.2% PVA followed by solvent evaporation on a stir plate (400 rpm for 3h). Microspheres were harvested by centrifugation (2000 rcf), washedthree times in ddH₂O, lyophilized to a dry powder, and stored at 4° C.

For each anti-VEGF agent, encapsulation efficiency (EE) was determinedfrom the radioactivity measured using a gamma counter (Cobra-IIAuto-Gamma, Packard Instrument Co., Meriden, Conn.) before and aftermicrosphere preparation. EE was defined as the percent-drug within themicrospheres relative to the theoretical loading amount.

To measure microsphere diameter, microspheres were fabricated asdescribed above with non-radiolabeled drug. A sample was suspended inddH₂O and imaged with a microscope (×20 objective, Carl Zeiss, Germany).Images of microspheres (n=135) were analyzed using ImageJ (developed byWayne Rasband, National Institutes of Health, Bethesda, Md.) to measurethe diameter.

Morphology of the microspheres was also examined by scanning electronmicroscopy (SEM). Dry microspheres were mounted on double-faced adhesivecarbon tape on metal stubs, sputter-coated with gold (E5000M S.E.M.Coater, Polaron Equipment Ltd., UK), and further analyzed in a scanningelectron microscope (JSM-5900LV, Jeol USA, Inc., Peabody, Mass.).

PNIPAAm-PEG-diacrylate (DA) hydrogels were synthesized using a methoddescribed in Osswald C. R. et al. Hydrogels were made in triplicate andprepared by dissolving PEG-DA (MW 575 Da, 2 mM), N-tert-butylacrylamide(47 mM), and ammonium persulfate (13 mM) in 1× Dulbecco's PBS (withCaCl₂ and MgCl₂). Then, NIPAAm (350 mM) was added to create the hydrogelprecursor. Microspheres (15 mg) were suspended in 1 mL of precursor in a2 cc microcentrifuge tube and kept on ice. Polymerization of thehydrogel was initiated by adding N,N,N′,N′-Tetramethylethylenedi amine(168 mM). After polymerization, hydrogels were collected and washedthree times in ddH2O.

Release of anti-VEGF agents was determined using a separation methodalso described in Osswald C. R. et al. Briefly, 15 mg of microspheresand 1 mL of the DDS (15 mg microspheres in 1 mL hydrogel) were suspendedin 1.5 mL of 1×PBS to ensure detectable levels of anti-VEGFconcentration throughout release. Release profiles were conducted at 37°C. under mild agitation and at predetermined intervals, 1 mL ofsupernatant was removed after a brief centrifugation and replaced withan equal volume of fresh buffer. Supernatants were read using a gammacounter (Packard). Cumulative release was calculated as a percent ofencapsulated drug and was considered complete when the microspheres hadcompletely degraded based on visual inspection. The D3 was defined asdrug released within the first 24 hours.

For the bioactivity studies, non-radiolabeled release samples were keptat −80° C. until bioactivity testing. To determine bioactivity, humanumbilical vascular endothelial cells (HUVECs) were cultured,trypsinized, and seeded in 96-well plates at 5000 cells per well ingrowth medium (EGM-2, Lonza Inc., Allendale, N.J., USA). After the cellshad grown to −80% confluence (48 hr), cell growth was arrested bywashing the plates twice with 1×PBS and then adding 50 μl of basalgrowth media (EBM-2, Lonza) to each well. To achieve VEGF-inducedproliferation, 25 μL of 10 ng/mL exogenous VEGF (taken from a BulletKit,Lonza) and 25 μL of a release sample of anti-VEGF were then added toeach well. Each release sample was done in duplicate for a total of sixsamples per time point (microspheres alone and DDS made in triplicate).A solution containing a tetrazolium compound[3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium,inner salt; MTS] and an electron coupling reagent (phenazineethosulfate) (MTS assay, CellTiter 96® AQ_(ueous) One Solution Reagent;Promega, Madison, Wis., USA) was used according to the manufacturer'sprotocol to determine cell proliferation and cytotoxicity after two daysof exposure to the release samples. In the MTS assay, cell proliferationis proportional to the optical density (OD) of the sample when readusing a spectrophotometer at 490 nm from 1-4 hours. Cell proliferationwas normalized relative to those wells receiving only exogenous 10 ng/mLVEGF (negative control). Positive controls consisted of 25 μL of aclinical dose of drug (10 mg/mL ranibizumab or 40 mg/mL aflibercept)added to wells containing 10 ng/mL VEGF.

All values are reported as mean±standard error and in all graphs, errorbars represent standard error. Significant differences in microspherediameter, EE, and D3 were determined using Student's t-test. For thebioactivity studies, significance at each time point was determinedusing Student's t-test. Significance between the test conditions wasdetermined using one-way ANOVA followed by the Holm-Sidak test. Unlessotherwise noted, significance represents p<0.05.

PLGA microspheres loaded with either ranibizumab or aflibercept weresuccessfully created and suspended within a thermo-responsivePNIPAAm-based hydrogel. Representative images of ranibizumab- andaflibercept-loaded microspheres can be seen in FIGS. 1A and 1C,respectively, with microspheres appearing smooth and non-porous. Afterthe microspheres were suspended within the hydrogel, the DDS remainedinjectable through a 28 G needle at room temperature (22° C.; FIGS. 1Band 1D). Particle size and EE for ranibizumab- and aflibercept-loadedmicrospheres alone and suspended in hydrogel can be seen in Table 1. TheEE was not significantly different for either drug when comparingmicrospheres alone to those in hydrogel (p>0.65); however,ranibizumab-loaded microspheres had twice the EE compared toaflibercept-loaded microspheres.

TABLE 1 Characteristics of release for both ranibizumab- andaflibercept-loaded microspheres and the same suspended in hydrogel.Encapsulation Efficiency Initial Burst Diameter Percent Weight PercentWeight (μm) (%) (μm) (%) (μm) Ranibizumab Alone 7.5 ± 0.4 91.3 ± 2.5149.6 ± 3.7 50.3 ± 2.1 113.2 ± 4.7 in 89.5 ± 2.8 106.2 ± 3.0 21.0 ± 2.0 22.2 ± 2.2 hydrogel Aflibercept Alone 8.0 ± 0.3 44.8 ± 2.0 169.8 ± 3.483.3 ± 2.1  65.3 ± 0.7 in 44.6 ± 1.4  65.4 ± 0.9 20.1 ± 0.8  13.1 ± 0.5hydrogel

The hydrogel provided a significant reduction in the D3 for bothanti-VEGF drugs. For ranibizumab, the D3 was reduced by 58% (p=0.0003)and for aflibercept, the D3 was reduced by 76% (p=0.0002) (Table 1).Beyond the first week of release, release from microspheres alone forboth ranibizumab and aflibercept was minimal (FIGS. 2A and 3A). Thisrelease phenomenon is often seen in protein-loaded microsphere. Itresults from the majority of protein being surface-bound, which quicklydiffuses away. Thus, there is minimal protein entrapped within the bulkof the microsphere, which precludes release at later time points frommicrospheres alone.

In contrast, suspending microspheres in hydrogel resulted in steadyrelease of 0.153 μg/day of ranibizumab (FIG. 2B) and 0.065 μg/day ofaflibercept (FIG. 3B) from Days 7-196. Thus, the hydrogel providessufficient resistance to the diffusion of surface-bound anti-VEGF, whichyields controlled release throughout. Additionally, both ranibizumab-and aflibercept-loaded microspheres completely degraded after 154 days,whereas suspending the microspheres in hydrogel extended the release ofboth drugs by 27.2% to 196 days.

Hydrogels were checked for translucency beginning on Day 154 (time atwhich microspheres alone had completely degraded but not microspheres inhydrogel) to determine whether the microspheres had degraded. This wasdone by lowering the temperature of the hydrogel from 37° C. to 22° C.and determining, by visual inspection, whether microspheres remainedwithin the hydrogel. The hydrogels returned to their translucent stateat 203 days; thus, degradation of the microspheres was determined to becomplete by 196 days. (11) An additional 25.1±12.1% (27.3 μg) ofranibizumab and 32.7±1.1% (21.4 μg) of aflibercept remained entrapped inthe hydrogel after complete degradation of the microspheres (FIGS. 2 and3, respectively; circled).

Bioactivity was maintained throughout release. A proportional increasein HUVEC proliferation was observed with increased VEGF concentration(FIGS. 4 and 5, left panels). Positive controls (clinical dose of eachdrug) significantly inhibited VEGF-induced proliferation of HUVECs(p<0.0007) and no difference was seen between positive controls andbasal growth medium (BGM, p>0.92).

For both drugs, significant inhibition of VEGF-induced HUVECproliferation was observed during the D3 (p<0.05). At all subsequenttime points, HUVEC proliferation was less for the drug-loaded DDScompared to the non-loaded DDS counterparts (labeled as “Blank-gel” inFIGS. 4 and 5). For both the ranibizumab- (FIG. 4, right panel) andaflibercept-loaded DDS (FIG. 5, right panel), significant inhibition ofHUVEC proliferation was observed at many time points throughout release,including later release time points (p<0.05). Additionally, both drugsremain bioactive throughout release as indicated by a lower normalizedOD for drug-treated HUVECs compared to control at every time point(FIGS. 4 and 5, right panes). However, at various time points, aninsufficient amount of bioactive anti-VEGF was available tosignificantly inhibit HUVEC proliferation (p>0.05).

Controlled and extended release of bioactive ranibizumab and afliberceptwas achieved for ˜200 days by encapsulating these drugs in PLGA 75:25microspheres suspended in a PNIPAAm-based hydrogel. Compared tomicrospheres alone, the hydrogel significantly lowered the D3 of bothdrugs and extended the release of each by almost a third. Additionally,the hydrogel allowed for significant bioactivity of both drugs to bemaintained throughout. If these results are verified in vivo, bioactiveand controlled release for nearly 200 days is a significant step towardsreplacing the current monthly/bimonthly treatment regimens approved forthese drugs.

Interestingly, ranibizumab-loaded microspheres had twice the EE ofaflibercept-loaded microspheres even though the theoretical loading wasthe same and the drugs were treated identically in the fabrication ofthe microspheres. A likely cause for this is the difference in MWbetween the drugs. It has been shown across a variety of fabricationtechniques, including the double-emulsion technique used in this study,that smaller proteins tend to have a higher EE. Larger proteins tend toremain surface-bound at the polymer-water interface, as was evidenced bythe significantly higher D3 for aflibercept (MW 115 kDa) compared toranibizumab (MW 48 kDa). Additionally, charge and tensioactiveproperties of the proteins may have contributed to the differences seenin EE.

Suspending microspheres within the thermo-responsive hydrogel accordingto this invention significantly reduced the D3 by 58% and 76% forranibizumab and aflibercept, respectively. The reduction in D3 may beattributed to the fabrication of hydrogel (e.g., surface diffusion ofdrug into hydrogel precursor prior to gelation) and the hydrogel actingas a diffusion barrier. A larger protein would see a greater diffusionbarrier; thus, it was anticipated and shown that the reduction in D3 wasgreater for aflibercept than ranibizumab. Importantly, suspending themicrospheres within the thermo-responsive hydrogel yielded steadier andmore controlled release compared to microspheres alone. Thismicrosphere-hydrogel DDS approach has been shown previously to providebetter control and extended release compared to microspheres alone.Thus, the DDS of this invention has two potential benefits for thetreatment of posterior segment eye diseases: it would prevent theindependent movement of microspheres within the eye as well as provideprolonged and controlled release of anti-VEGFs.

For both anti-VEGFs, a residual amount of drug was observed trapped inthe hydrogel at the end of the study (final time point, FIGS. 2 and 3).In a study by one of the inventors using PNIPAAm-based hydrogel toencapsulate BSA and immunoglobulin G (IgG), a similar entrapment wasobserved. Aflibercept had a higher residual entrapment (32.7±1.1%) thanthe smaller ranibizumab (25.1±12.1%), which had a higher residualentrapment than the smaller ovalbumin (45 kDa) observed in our priorstudy (8.01±0.76%). Non-degradable hydrogels were used intentionally inthis study to simplify the experimental factors and focus on the releaseof ranibizumab and aflibercept. The issue of entrapped drug using thehydrogel can be mediated by utilizing fully biodegradable hydrogels.Hydrogels with varying lengths of degradation can have an effect onrelease from microspheres. For example, FIG. 6 summarizes data from anexperiment (though 180 days) showing release of ranibizumab frombiodegradable microspheres and a biodegradable thermo-responsivehydrogel. Different amounts of crosslinker (PLLA-PEG-DA) were used tochange the degradation rate.

Including the residual entrapped drug, incomplete release (that is, notachieving 100% cumulative release) was still observed. This phenomenonis frequently reported in many protein-loaded PLGA microsphere deliveryplatforms and may be due to protein instability. Protein aggregation andnon-specific protein adsorption are typically the leading causes ofincomplete release and differences in the isoelectric point (pI) of eachdrug may be to blame. At pH lower than the pI, proteins carry a netpositive charge and vice versa. In a previous study using ovalbumin,complete release was observed. The pI of ovalbumin is ˜4.6, whereasmonoclonal antibodies (e.g., ranibizumab) typically have a pI of ˜8.Thus, performing these experiments at physiological pH (7.4) may haveinfluenced the release due to possible charge interactions between thedrugs, PLGA, and hydrogel.

Significant bioactivity is seen during late release time points (i.e.,beyond 100 days), a timeframe well beyond the D3 where a large amount ofdrug was released. In our initial characterization of this DDS withovalbumin, a large second burst occurred after Day 70 and release beyondthe second burst occurred at a greater rate than prior to the secondburst. The D3 can be readily explained by diffusion of drug located nearthe particle surface through short diffusion pathways. As PLGA swellsand degrades, the pores become large enough to allow for the release ofentrapped protein, which can lead to a second burst. Although a largesecond burst was not observed in this study for either anti-VEGF-loadedDDS, it is likely that at these later time points, even though theamount of protein was small, the protein was likely intact and thus asbioactive as during the D3. Proteins at interfaces (e.g., surface-boundprotein during the IB) tend to easily denature whereas proteinsentrapped within the microsphere (i.e., release at later time points)tend to remain intact and bioactive.

The release of bioactive anti-VEGF for over six months is a significantstep toward eliminating monthly or bimonthly bolus injections of drugsneeded to treat many posterior segment eye diseases. The extended andcontrolled release of anti-VEGF agents that the DDS of this inventiondemonstrates may be a significant advancement in the delivery of proteintherapeutics to the eye once validated in vivo. The DDS of thisinvention has the potential to improve upon the socio-economic impactassociated with recurrent injections and to lower the risk forassociated potential complications.

To further validate the potential of the DDS of this invention to treatposterior segment eye diseases, the in vivo efficacy of the DDS wasdetermined using a rat model of CNV. Animal models of CNV fall intothree categories: laser and light induced, surgically induced, andtransgenic and knockout mouse models. These models have a commonbeginning: a break or defect in Bruch's membrane. Development of CNV isa dynamic process with initiation, maintenance, and involution stages.Currently, the most common way to induce CNV in animals is through theuse of laser photocoagulation as it is relatively simple to create, isinexpensive, and is reproducible.

In rats, approximately 75% of lesions are vascularized within the firstweek following CNV induction; after ten days, CNV is fully developed andremains so for at least 35 days after induction. Additionally, it hasbeen found that the highest rate of blood vessel growth (3-10 days afterinduction) correlates with the peak expression of VEGF and its receptor.Furthermore, using vascular cast images it has been shown that featuresof laser-induced CNV in rats were similar at one and three months butthat CNV had atrophied by six months. Accordingly, lesions should bemonitored at least weekly in the beginning of CNV development, with theassumption that CNV is not fully developed until two weekspost-induction and lesions may begin to atrophy after three months.

Several methods exist to quantify CNV lesion growth. Often, a four-tiergrading scale is used ranging from no leakage to severeleakage/hyperfluorescence, where FA images are analyzed by trainedretinal specialists, averaged, and compared. However, this methodremains somewhat subjective and can be influenced by image properties(e.g., brightness and contrast). More complex and objective methodsinvolve flat-mount preparations, immunohistochemistry with serialreproductions, and OCT. However, these complex methods can quicklybecome quite expensive and time consuming. To address the issues ofobjectivity, ease of use, and cost-effectiveness, a novel technique wasdeveloped based on FA images and Otsu thresholding to monitor CNV lesiongrowth.

Otsu thresholding allows for the automatic reduction of a grayscaleimage into a binary foreground/background image such that an optimumthreshold separating the two classes maximizes the inter-class varianceof pixel intensity. This can be extended to separate intermediate levelsand has been implemented as a plug-in for the open-source softwareImageJ (NIH). To quantify CNV areas, three levels (or “regions”) aredefined: background, diffuse leakage, and CNV (FIG. 7). When determiningCNV areas, the CNV lesion is outlined with blood vessels avoided. Thelesion area is measured in pixels, which can then be converted into areausing a scaling factor.

The ability of controlled and extended release of anti-VEGF from the DDSto inhibit CNV lesion growth was tested and compared to that of acorresponding bolus injection of anti-VEGF. Six treatment groups wereobserved: non-treatment, drug-free DDS, ranibizumab-loaded DDS,ranibizumab bolus, aflibercept-loaded DDS, and aflibercept bolus. Alongwith determining CNV areas, ocular health was monitored to demonstratethat the DDS was biocompatible and well-tolerated in vivo. It washypothesized that controlled and extended release of anti-VEGF from theDDS will more effectively treats CNV than a bolus injection and thatocular health will be maintained throughout the study.

All animal procedures were in accordance with protocols approved by theInstitutional Animal Care and Use Committee at the Illinois Institute ofTechnology, and with the principles embodied in the statement on the useof animals in ophthalmic and vision research adopted by the Associationfor Research in Vision and Ophthalmology. Long-Evans male rats (1-3months, 300-350 g) were purchased from Harlan Laboratories(Indianapolis, Ind., USA). Animals were anesthetized using 80 mg/kg ofketamine hydrochloride (Fort Dodge Animal Health, Fort Dodge, Iowa, USA)and 10 mg/kg xylazine (AnaSed® Injection, Akorn, Inc., Decatur, Ill.,USA) via intraperitoneal (IP) injection. Proparacaine drops (Bausch andLomb, Rochester, N.Y., USA) were used to anesthetize the corneasthroughout the procedure and pupils were dilated using phenylephrine(Bausch and Lomb) and atropine drops (Bausch and Lomb). Heart rate andblood oxygen saturation were monitored with a PulseOximeter (8500AV;Nonin Medical Inc., Plymouth, Minn., USA). Animals were placed on acustom-built heated stage and monitored to maintain a core bodytemperature of 37° C.

Non-radiolabelled, ranibizumab-, and aflibercept-loaded microsphereswere prepared. Fifteen (15) mg of microspheres were put in a 2 ccmicrocentrifuge tube and placed under UV light for 30 min to sterilize.Under sterile conditions, the hydrogel precursor and initiator weresterile-filtered using 13 mm syringe filter (0.22 μm, Fisherbrand,Thermo Fisher Scientific, Waltham, Mass., USA). One (1) mL of hydrogelprecursor was added to the microspheres and the tube was invertedseveral times to suspend the microspheres. The initiator was then added,and free radical polymerization occurred on ice for 30 min. Hydrogelswere washed three times in sterile PBS, loaded into 0.5 cc U-100 insulinsyringes (28G½; Becton Dickinson & Co., Franklin Lakes, N.J., USA), andstored at 4° C.

Laser photocoagulation was performed using an argon-green laser(AKC-8000, NIDEK, Inc., Fremont, Calif., USA) attached to a slit lampwith a laser power of 400 mW, duration of 100 ms, and spot diameter of50 μm. Using a 90-diopter lens, the posterior pole of the eye was viewedand the laser beam focused on the retina. Five to six lesions per eyewere induced two to three disc diameters from, and centered on, theoptic disc. Laser-induced disruption of Bruch's membrane was identifiedby the appearance of a bubble at the site of photocoagulation. Laserspots that did not result in the formation of a bubble were excludedfrom study.

Animals were separated into six groups to determine the ability of theDDS of this invention to treat CNV. For analysis purposes, lesions wereconsidered independent within each treatment group and each groupcontained two rats for a total of four eyes and up to 24 lesions pertreatment group. The treatment groups were as follows: 1) control group,which did not receive any treatment or injection after laser induction(“non-treatment”); 2) non-loaded DDS (“Blank-gel”); 3)ranibizumab-loaded DDS (“Rani.-gel”); a single bolus injection ofranibizumab at clinical dose (10 mg/mL, “Rani.”); a single bolusinjection of aflibercept-loaded DDS (“Aflib.-gel”); aflibercept atclinical dose (40 mg/mL, “Aflib.”). All IVT injections were 5 μL andperformed immediately after CNV induction.

The corneal electroretinogram (ERG) is a non-invasive measurement thatrepresents the overall electrical activity of the retina in response toa stimulus flash. It is routinely used to assess retinal cellularactivity and in toxicology studies. Each component of the ERG can beused to assess different retinal cell types and any alterations in thefunction of the retina due to the DDS would appear as changes inamplitude or sensitivity of various ERG components. Under scotopicconditions, the ERG has two important components: the a-wave and b-wave.The rising edge of the negative a-wave is generated by thephotoreceptors in the outer retina; the positive b-wave is generatedprimarily by the bipolar cells and/or Müller cells. Another component ofthe ERG is the c-wave, which originates in the RPE and can be used toassess the functional integrity of the photoreceptors, the RPE cells,and the interactions between them; however, for the purposes of thisresearch, the c-wave was not recorded or analyzed.

Under dim red light prior to and at each time point after CNV inductionand DDS injection, ERG experiments were performed under scotopicconditions, with animals dark-adapted overnight prior to the experiment.Electroretinograms, in response to the full-field Ganzfeld stimulation,were recorded by a gold wire loop placed on the cornea. The referenceand ground electrodes were 30 G platinum subdermal needle electrodes(SAFELEAD™ F-E2, Grass Products, Natus Neurology, Warwick, R.I., USA)inserted into the cheek and nape of the neck, respectively. A-wave andb-wave intensity responses were recorded by presenting single flashes ofincreasing intensity (5×10⁻⁴ to 305.7 sc cd·s·m⁻²) and allowing the eyeto dark adapt for 1 min after each flash. The intensity-response valuesfor the a-wave and b-wave were fit with the Naka-Rushton equation, todetermine a half-saturation intensity of the responses and maximalresponse.

Prior to treatment, post-treatment, and weekly thereafter, intraocularpressure (IOP) measurements were taken using an applanation tonometer(TONO-PEN® XL, Medtronic, Minneapolis, Minn., USA). The tono-pen hasbeen shown to effectively measure rat TOP. It should be noted thatgeneral anesthesia, including combined ketamine and xylazine, has beenshown to induce rapid and substantial decreases in IOP as well asincreased inter-animal variability in IOPs.

A confocal scanning laser ophthalmoscope (cSLO) system (HeidelbergRetina Angiograph (HRA), Heidelberg Engineering, Heidelberg, Germany)was used to image the retina. FA still images were captured at 1-2 min,10 min and 20 min after IP injection of 0.5 mL of 20% fluorescein dye(Sigma Aldrich, St. Louis, Mo., USA) and were used for quantification ofCNV lesion areas. All data was acquired prior to treatment and at 1, 2,4, 8, and 12 weeks post-injection.

CNV lesion areas were quantified using late-phase FA images. The use oflate-phase (20 min) images allows time for the fluorescein to clear thenormal vasculature, preventing vessel hyperfluorescence from interferingwith area measurements. Additionally, only images that were centered onthe lesion, well-focused on the choroid, non-saturated, and had evenillumination were used to quantify CNV.

An example of image processing using ImageJ is presented in FIG. 7. Themulti-Otsu thresholding (MOT) plugin for ImageJ was set to three levels,or “regions,” defined as: background (FIG. 7B), diffuse leakage(non-hyperfluorescent leakage; FIG. 7C), and CNV (hyperfluorescentleakage; FIG. 7D). The threshold value within ImageJ is then adjusted toinclude all pixels within the “Region 2” (CNV) plug-in output. The CNVlesion was then outlined by the user, being careful to avoid vessels andother non-lesion pixels (FIG. 7E). The measured number of pixels wasthen converted into area using a scaling factor based on the opticalproperties of the rat eye and field of view of the image using theequation:

${SF} = {\frac{\pi}{180} \times N^{\prime}F^{\prime} \times \frac{FOV}{RES}}$

where SF is the scaling factor in square micrometers per pixel, N′F′ isthe distance from the second nodal point of the eye to the retina inmicrometers, FOV is the field of view in degrees, and RES is the lengthof the image in pixels. For the Long-Evans rats used in this study, SFwas 3.35 μm². Note that within ImageJ, the MOT plugin requires an 8-bitgrayscale image and measurements must be set to “Limit to threshold”under the “Analyze >Set measurements . . . ” tab, otherwise all pixelswithin the outline will be counted rather than only within the lesion.

The parameter values obtained from the Naka-Rushton analyses and IOPmeasurements were compared using the paired t-test verses control. Todetermine significant differences in CNV lesion areas, one-way ANOVA wasperformed and all pairwise multiple comparisons were done using theHolm-Sidak method. Unless otherwise noted, significance representsp<0.05. All values are reported as sample mean±standard error of themean.

In a previous, thorough evaluation of non-drug-loaded thermo-responsivehydrogel, no adverse effects were observed out to four weeks post-IVTinjection. As the microsphere-hydrogel DDS was to be evaluated over thecourse of 12 weeks, an additional ocular health and safety evaluationwas performed including ERG analysis and IOP measurements.

The outer retina was evaluated by the a-wave, which reflects theactivity of the photoreceptors. The sensitivity of the a-wave wasexamined through comparison of parameters obtained from the Naka-Rushtonanalyses and the intensity-response of the a-wave was determined before(control) and after injection of the DDS and the bolus drugcounterparts. As was expected, no significant differences were seen inmaximal a-wave response (control: 543.1±51.7 mV) or half-saturationintensity (control: 2.9±1.6 sc cd·s·m⁻²) of non-treated animals (FIG. 8,top). Additionally, no significant differences in a-wave parameters wereseen for drug-free DDS treated animals (control Amax: 400.9±23.7 mV,control σA: 2.6±0.8 sc cd·s·m⁻²; FIG. 9, top). For both theranibizumab-loaded DDS (control Amax: 464.1±19.9 mV, control σA: 2.2±0.5sc cd·s·m⁻²) and bolus ranibizumab (control Amax: 419.9±24.0 mV, controlσA: 2.6±0.8 sc cd·s·m⁻²) treated animals, no significant differenceswere seen in a-wave parameters (FIGS. 10 and 11, top; respectively).However, for both the aflibercept-loaded DDS and bolus aflibercepttreated animals, a significant difference in maximal a-wave response isseen between control and Week 1-Week 12 (p<0.05). For theaflibercept-loaded DDS treated animals, the maximal a-wave responsepost-treatment was ˜20% lower than control (Amax: 516.0 51.1 mV) at alltime-points post-treatment (FIG. 12, top) and for the bolus aflibercepttreated animals, it was ˜40% (control Amax: 616.4±25.4 mV; FIG. 13,top). No significant difference in half-saturation intensity wasobserved for either aflibercept-loaded DDS (control σA: 4.5±2.5 sccd·s·m⁻²) or bolus aflibercept (control σA: 4.6±0.9 sc cd·s·m⁻²) treatedanimals.

Inner retinal function was evaluated by the b-wave, which reflects theactivity of the bipolar cells and/or Müller cells. As with the a-wave,the sensitivity of the b-wave was examined through comparison ofparameters obtained from the Naka-Rushton analyses and theintensity-response of the b-wave was determined before (control) andafter injection of the DDS and the bolus drug counterparts. As expected,no significant differences were seen in maximal b-wave response(control: 748.5±70.7 mV) or half-saturation intensity (control GB:2.6×10-3±8.0×10-4 sc cd·s·m⁻²) of non-treated animals (FIG. 8, bottom).A significant increase of −35% in maximal b-wave response was seen onlyat Week 12 compared to control for drug-free DDS treated animals(control B max: 684.7±37.8 mV; FIG. 9, bottom; p<0.05), suggesting thatthe DDS may cause a greater depolarization of the ON-bipolar cells ofthe retina in the long-term. Similar to the a-wave, for both theranibizumab-loaded DDS (control Bmax: 987.9±91.0 mV, control GB:3.3×10⁻³±1.0×10⁻³ sc cd·s·m⁻²) and bolus ranibizumab treated animals(control Bmax: 734.3±65.0 mV, control GB: 4.1×10⁻³±1.1×10⁻³ sccd·s·m⁻²), no significant differences were seen in b-wave parametersafter treatment (FIGS. 10 and 11, bottom; respectively). For theaflibercept-loaded DDS treated animals, no significant differences inb-wave parameters were observed (control Bmax: 790.4±51.8 mV, controlGB: 3.0×10⁻³±6.0×10⁻⁴ sc cd·s·m⁻²; FIG. 12, bottom). Surprisingly, forthe bolus aflibercept treated animals, a significant decrease in maximalb-wave response is seen post-treatment through Week 12, with the maximalb-wave response being ˜45% lower than controls (control Bmax:1236.1±121.5 mV, control GB: 4.6×10⁻³±1.4×10⁻³ sc cd·s·m⁻²; FIG. 13,bottom). Although a significant decrease in maximal b-wave response wasseen in these animals after treatment, control maximal a- and b-waveresponses for these animals were >50% higher than all other controlmeasurements, which may suggest abnormal experimental conditions duringcontrol ERG measurements (e.g., residual eye drops enhancing the contactof the cornea electrode or the position of the eye being closer to theflash stimulus than normal).

IOP measurements were recorded prior to and immediately after IVTinjection and at every time point thereafter. Table 2 shows the averageIOP measurements for all treatment groups at each time point. Theaverage control IOP values for all animals was 18.8±0.1 mmHg. As wasanticipated, a significant increase of ˜37% in IOP is seen immediatelyafter IVT injection in all treated animals (p<0.05). Increased IOPremained at Week 1 for the aflibercept-loaded DDS and both bolus treatedanimals (23.3±1.0 and 25.0±1.8 mmHg, respectively). By Week 2, onlybolus aflibercept treated animals had increased IOP (25.0±1.8 mmHg).Interestingly, at Weeks 8-12, the ranibizumab-loaded DDS treated animalshad significantly lower IOP verses control (17.0±0.8 and 17.3±1.0 mmHg,respectively) and at Week 12, bolus ranibizumab treated animals also hadsignificantly decreased IOP (17.3±1.5 mmHg). It should be noted that allIOP values were well within the range of what is normal in rats (15-25mmHg).

TABLE 2 Average intraocular pressure measurements for all treatmentgroups Post- Treatment Control Injection Week 1 Week 2 Week 4 Week 8Week 12 Non- 17.3 ± 2.1 — 22.0 ± 4.0  19.3 ± 0.6  19.0 ± 0.0 19.7 ± 0.6 17.3 ± 1.2  treatment Drug-free 19.0 ± 2.4 25.5 ± 4.4* 21.8 ± 1.3  21.8± 1.5  21.3 ± 1.0 19.0 ± 0.8  18.8 ± 1.5  DDS Rani.-gel 19.5 ± 1.3 27.0± 0.8* 21.5 ± 1.7  20.8 ± 1.3  20.8 ± 1.0 17.0 ± 0.8* 17.3 ± 1.0*Rani.-bolus 20.0 ± 2.2 26.5 ± 1.3* 24.3 ± 1.7  20.3 ± 2.2  20.0 ± 2.019.8 ± 1.7  17.3 ± 1.5* Aflib.-gel 18.5 ± 1.3 25.0 ± 3.6* 23.3 ± 1.0*20.8 ± 1.0  18.5 ± 2.1 19.8 ± 1.0  18.3 ± 1.3  Aflib.-bolus 18.0 ± 1.426.5 ± 1.7* 25.0 ± 1.8* 20.8 ± 1.0* 18.2 ± 1.3 19.3 ± 2.2  17.8 ± 1.5 *Significant differences in IOP compared to each treatments' controlmeasurement (p < 0.05)

The MOT technique effectively delineated the CNV lesion from thesurrounding retina in the experimental laser model, allowing forobjective quantification of CNV lesion areas. The CNV lesion areas forall treatment groups can be seen in FIG. 14.

Representative composite FA images show that the severity of CNV lesionsat Week 12 for each treatment group was approximately the same (FIG.15). Although CNV is assumed to not be fully developed at Week 1, it hasbeen presented for completeness. As anticipated, no significantdifference was observed between non-treated animals and drug-free DDStreated animals at any time point. From Week 4-Week 12, theranibizumab-loaded DDS had significantly smaller lesion areas thannon-treated animals and by Week 12, these treated animals had CNV lesionareas that were 60% smaller (p<0.0002). Additionally, beginning on Week1, ranibizumab-loaded DDS treated animals exhibited smaller CNV lesionareas than bolus ranibizumab treated animals, however this differencewas not significant. Furthermore, bolus ranibizumab treated animals onlyhad a significant decrease in lesion area at Week 12. Bolus aflibercepttreated animals had significantly smaller CNV lesions compared tonon-treated animals and aflibercept-loaded DDS at Week 1, however it isunknown whether this difference is meaningful as CNV growth is not fullydeveloped until ˜10 days. Additionally, bolus aflibercept treatedanimals had significantly smaller CNV lesions compared to non-treatedanimals at Week 4 and Week 12. However, by Week 8, aflibercept-loadedDDS treated animals had significantly smaller CNV lesions than bolus andDDS treated animals from Week 8-Week 12 (p<0.05).

For both anti-VEGF-loaded DDSs, it is important to note that significantdecreases in CNV lesion growth were achieved that exceeded that of bolusadministration in the long term. With a 5 μL bolus injection of aclinical dose of either anti-VEGF, 50 μg of ranibizumab or 200 μg ofaflibercept was delivered. In comparison, over the course of 12 weeksonly ˜450 ng of ranibizumab or ˜200 ng aflibercept was delivered fromthe DDS of this invention. Thus, there is great potential of controlledand extended release from the DDS of this invention in the treatment ofCNV.

The efficacy of the DDS to treat laser-induced CNV in rats was evaluatedin addition to determining the safety and biocompatibility of the DDS.Retinal cellular function was evaluated using ERG parameters determinedby fitting the intensity-response curve to the Naka-Rushton equation.Intraocular pressure measurements were also taken. Finally, CNV lesionareas were measured at every time point to determine whether the DDSeffectively treated laser-induced CNV in Long-Evans rats.

In animals treated with the drug-free DDS, photoreceptor cellularfunction was maintained and normal ERG activity was observed throughoutthe 12 weeks post-treatment. This result expands the time frame forwhich the thermo-responsive hydrogel DDS has been shown to cause nolong-term physiological effects on the outer retina. In a previousstudy, the thermo-responsive hydrogel was evaluated for four weeks andit did not cause any long-term physiological effects on thephotoreceptors. The lack of change in the outer retina was notsurprising as the photoreceptor cells are located farthest from thehydrogel injection. It follows, then, that the inner retinal cells aremore likely to experience any adverse effects due to the hydrogelinjection, and changes in inner retinal function would be apparent inthe b-wave. As was shown, inner retinal health and activity wasmaintained through Week 8. Interestingly, a ˜35% increase in maximalb-wave response was observed at Week 12 in all animals, which couldimply a meaningful adverse effect in long-term biocompatibility.However, no significant changes in a- or b-wave half-saturationintensity, which indicates the sensitivity of the retina toillumination, were seen in any treatment group throughout the entirestudy. The human rod system requires an intensity change of greater than14% for an observer to perceive a change in light intensity. Thus, whileno change in half-maximal sensitivity was observed, it is possible thatthe large increase in maximal b-wave response observed causedperceptible changes in the rodent's vision whereby the rat was moresensitive to light at Week 12 compared to control.

For both the bolus ranibizumab and ranibizumab-loaded DDS treatedanimals, no significant differences were seen in either maximal a- andb-wave response or half-maximal a- and b-wave response, furtherindicating that the DDS does not induce adverse effects on retinalactivity. Although no changes were observed in maximal b-waveparameters, a small by significant decrease (˜10%) was observed inaflibercept-loaded DDS treated animals after treatment. For reasonsmentioned above, this slight decrease would likely not cause perceptiblechanges in the rodent's vision. More surprising was the large (˜40%),sustained decrease in maximal a- and b-wave responses for bolusaflibercept treated animals subsequent post-treatment. These bolustreated animals had a similar number and severity of lesions as theother treated groups (FIG. 14), which eliminates the possibility thatthe decrease in the full-field ERG maximal a- and b-wave responses wasdue to greater CNV area (i.e., decreased retinal activity). However,control measurements for these animals were >50% higher than all othercontrol measurements, which may suggest abnormal experimental conditionsduring control ERG measurements (e.g., residual eye drops enhancing thecontact of the cornea electrode or the position of the eye being closerto the flash stimulus than normal). In the future, histologicalevaluation of the enucleated eyes will confirm whether this sustaineddecrease in maximal a- and b-wave response was due to irreversible toxicretinal damage.

The clinical dose of aflibercept (50 μL injection of 40 mg/mLaflibercept) when injected into the human vitreous (3.8 mL) yields avitreal concentration of 0.52 mg/mL. In this study, 5 μL of 40 mg/mLaflibercept was injected into the rat vitreous (13.6 μL vitreousvolume), which translates into a vitreal concentration of 10.8 mg/mL,which is a much higher vitreal concentration than that used in humans.In a recent ex vivo study on the electrophysiological toxicity testingof aflibercept in bovine retina, a significant decrease (43.1%) in themaximal scotopic b-wave response was observed directly after exposure to0.1 mg/mL aflibercept. After washing away the aflibercept, the b-waveresponse was no longer significantly decreased. So, it is possible thatfor rats, the high bolus aflibercept concentration used in this studycaused irreversible toxic retinal damage. However, in a study inchinchilla rabbits, a clinical dose of aflibercept was administered andwith a vitreous volume of approximately 200-300 μL, this translates intoa vitreal concentration of 6.67-10 mg/mL. It was found that this highvitreal concentration did not alter any ERG parameters at 24 hr or 7days after treatment; however, rabbits were only dark-adapted for 30minutes prior to ERG recordings and this is an insufficient amount oftime for complete dark-adaptation to occur. Importantly, the decreasedmaximal a- and b-wave responses in bolus aflibercept treated animals wasmediated by using the DDS as aflibercept is delivered slowly over timeand at much lower concentrations (i.e., two orders of magnitude less)and no changes in ERG parameters were observed in aflibercept-loaded DDStreated animals.

Retinal function can also be affected by changes in IOP. In the currentstudy, there was a significant increase in comparison to control in IOPmeasurements immediately after the injection, which resolved by Week 2for all treatment groups. This indicates that the IVT injection of theDDS of this invention had minimal impact on IOP. Surprisingly, in thebolus ranibizumab and ranibizumab-loaded DDS treated animals, asignificant decrease in IOP was observed at later time points. However,all values obtained in this study are in agreement with the normal rangeof IOPs for Long-Evans rodents. Although the adult rat vitreous volumeis −14 μL, IVT injection volumes of up to 5 μL have been shown by Dureauet al. (2001) to have good reproducibility with minimal loss of injectedsolution. Therefore, any changes seen in the ERG parameters were likelynot the result of altered TOP. Based on the ERG parameter analysis ofthe drug-free DDS and the IOP values for all treatment groups, it isbelieved that the DDS is safe, biocompatible, and well-tolerated.

No significant differences were seen between non-treated and drug-freeDDS animals at any time point. This suggests that the significantdecreases in CNV lesion areas observed in bolus or anti-VEGF-loaded DDStreated animals were in fact related to the anti-VEGF treatment. It hasbeen demonstrated that much less drug delivered constantly yields abetter treatment outcome than a bolus injection in the long term.Ranibizumab-loaded DDS treated animals exhibited a significant decreasein CNV lesion areas from Week 4-Week 12 compared to non-treated animalswhereas bolus ranibizumab treated animals only had a significantdecrease in lesion area at Week 12. Due to the short half-life ofranibizumab (t½=9 days), after 12 weeks, the ranibizumab that wasdelivered via bolus injection was effectively gone. It is important tonote that ranibizumab was administered immediately after CNV inductionand that a sufficiently high vitreal concentration of ranibizumab dosepresent during the highest rate of blood vessel growth (3-10 days afterinduction), which correlates with the peak expression of VEGF and itsreceptor. This may imply that the large initial bolus ranibizumab doseprevented CNV from fully forming and thus allowed CNV lesion areas tosignificantly decrease by Week 12.

Bolus aflibercept treated animals had significantly smaller CNV lesionscompared to non-treated animals at Week 4 and Week 12. However,aflibercept-loaded DDS treated animals had significantly smaller CNVlesions than bolus treated animals from Week 8-Week 12. Thus, theaflibercept-loaded DDS had greater efficacy in the long-term.Aflibercept is a more potent anti-VEGF than ranibizumab as demonstratedclinically by the need to administer ranibizumab more often thanaflibercept to achieve a similar treatment outcome. This may explain theearly advantage of bolus aflibercept compared to the DDS, yet extendeddelivery of aflibercept significantly reduced CNV lesion areas eventhough the rate of release for aflibercept from the DDS was 2.3× lessthan that of ranibizumab.

Interestingly, both anti-VEGF-loaded DDSs had roughly the same decreasedCNV lesion areas at Week 12, with aflibercept-loaded DDS areas only 2%larger than the ranibizumab-loaded DDS areas. In our preliminary studyon extended release of dexamethasone, a potent corticosteroid, from theDDS of this invention (nanospheres rather than microspheres), it hasbeen demonstrated that extended release of dexamethasone from the DDS ofthis invention greatly reduced CNV lesions areas compared to bolusadministration while delivering less than 5% of the amount of bolusdexamethasone. Together with the results presented here, this suggeststhat the efficacy of a variety of antiangiogenic agents could be greatlyimproved by controlled and extended release from the DDS of thisinvention.

There is debate as to whether bevacizumab, a humanized variant of theanti-human VEGF-A monoclonal antibody, and ranibizumab, a humanizedmonoclonal antibody and derivative of bevacizumab, are effective intreating murine neovascularization. In a study evaluating the efficacyand safety of intravitreal injections of bevacizumab and ranibizumab inthe treatment of CNV in a rat model, no therapeutic effect was observedout to 28 days over a range of doses both above and below the clinicaldose of each drug. However, in several studies on the effect of usingbevacizumab to treatment corneal neovascularization, humanized anti-VEGFwas found to be effective in the rat. It has been shown that systemic ortopical application of bevacizumab significantly inhibitedinflammation-induced angiogenesis in the cornea of mice and thatbevacizumab binds to mouse VEGF-A, which was determined by Western blot,enzyme-linked immunosorbent assay (ELISA), and plasmon resonance assay.Additionally, it has been demonstrated that topical bevacizumab limitedcorneal neovascularization in rats it has been found thatsub-conjunctival injection of bevacizumab can significantly inhibitcorneal angiogenesis. Furthermore, the results presented above clearlydemonstrate the efficacy of ranibizumab in treating murine CNV.

In summary, controlled and extended release of anti-VEGFs from the DDSwas shown to be more efficacious than bolus anti-VEGF treated animals ina laser-induced CNV murine model. Both anti-VEGF-loaded DDSs exhibitedvery similar efficacy at Week 12, suggesting that improved efficacy mayextend to other antiangiogenic agents. Additionally, the DDS was shownto be safe, biocompatible, and well-tolerated, with no long-term adverseeffects observed in drug-free DDS treated animals.

Example 2

Sustained drug delivery system (DDS) for anti-vascular endothelialgrowth factors such as aflibercept is in great demand for bettermanagement of chronic neovascular eye diseases. However, maintainingdrug stability and bioactivity during DDS fabrication and long-termrelease remains a big challenge. The purpose of this example was toinvestigate the effects of varying microsphere formulation on theaflibercept stability during fabrication and release frommicrosphere-hydrogel DDS.

The aflibercept was encapsulated into poly(lactic-co-glycolic acid)(PLGA) microspheres using double emulsion technique. Effects of organicsolvents (dichloromethane (DCM), triacetin, or ethyl acetate) and bovineserum albumin (BSA) contents (w/v %) (0%, 4%, 8%, 12%, 16%, or 20%) onaflibercept stability during primary emulsification was investigatedusing a simulated microencapsulation test. Stability of afliberceptafter simulated emulsification was measured using enzyme-linkedimmunosorbent assays (ELISA) to determine optimal combination of organicsolvent and BSA. Effects of various Mg(OH)₂ loadings relative to PLGA(w/w %) (0%, 3%, 6%, or 9%) on aflibercept stability during release fromDDS were also investigated using ELISA.

In each organic solvent group with different BSA contents, 8% BSAcontents generated the highest bioactive aflibercept recovery rate:92.16±6.35% (n=3) in triacetin; 91.46±3.90% (n=3) in ethyl acetate; and97.26±5.38% in DCM (n=3). Based on the results, 8% BSA with DCMcombination provided an optimal recovery rate. Addition of Mg(OH)₂ toorganic phase improve maintenance of aflibercept stability duringrelease timeframe from DDS. The stability of aflibercept after one-monthrelease from DDS with various Mg(OH)₂ loadings was determined asfollows: 4.66±2.56% for 0% Mg(OH)₂; 19.64±4.35% for 3% Mg(OH)₂;1.95±1.24% for 6% Mg(OH)₂; and 1.97±1.45% for 9% Mg(OH)₂. It was foundthat 3% Mg(OH)₂ produced highest aflibercept stability after one-monthrelease from PLGA microsphere-based DDS.

The data suggested that a combination of BSA and DCM protectsaflibercept from interfacial stress during primary emulsification.Addition of Mg(OH)₂ in organic phase helped improve afliberceptstability during release timeframe. Incorporating optimal ratio of BSAand Mg(OH)₂ may improve the long-term release of aflibercept from ourDDS.

Methods Effects of Organic Phase and BSA Contents on StabilizingAflibercept During Primary Emulsification.

Simulated primary emulsification of double emulsion encapsulationprocess was used here to determine optimal combination of organic phaseand BSA loadings. Formulated clinical aflibercept stock solution (40mg/ml) was used. The tests were performed in the absence of PLGA. 100μ1of aflibercept stock solution (40 mg/ml) was dissolved in PBS bearingvarious amounts of BSA, then incorporated into organic phase at a 1:5v/v water-to-organic phase ratio. Then in the first emulsification test,this mixture was stirred at 3200 rpm for 90 seconds to evaluatewater/organic solvent interface effect. Aflibercept was extracted fromthe organic phase by adding 4 ml of PBS, stirred for 2 min more, andthen centrifuged at 5000 rpm for 10 min to accelerate phase separation(total o/w ratio was 1:8.2). The aqueous phase was collected and used toevaluate the recovered amount and stability of aflibercept by ELISAassay. 0.5 ml PBS was used as control for comparing to various organicphases.

The organic solvents were DCM, triacetin, and ethyl acetate. BSAloadings (w/v %) were 0% BSA, 4% BSA, 8% BSA, 12% BSA, 16% BSA, and 20%BSA.

Optimal Mg(OH)₂ Amount (w/w %) for Stabilizing Aflibercept DuringRelease for One Month.

Based on the combination of organic phase and BSA loadings, the optimalloadings of Mg(OH)₂ were determined. The Mg(OH)₂ amounts relative toweight of PLGA (w/w %) investigated were 0% Mg(OH)₂, 3% Mg(OH)₂, 6%Mg(OH)₂, and 9% Mg(OH)₂. Release samples were collected at predeterminedtime points during first month release. Two different kinds of ELISAassays were designed and used to measure total protein (Total ELISA) andbioactive protein (Activity ELISA), respectively, to study proteinstability. Anti-human immunoglobulin coated strips were used for TotalELISA, whereas VEGF165 coated strips made Activity ELISA.

Results Combination of Organic Phase and BSA Loadings for StabilizingAflibercept During Primary Emulsification

FIG. 16 represents the recovery rates of the table below. For Triacetinthere were no significant differences among various BSA loadings. ForEthyl Acetate, 20% BSA loading gave significantly lower afliberceptrecovery rate. For DCM, 20% BSA loading gave significantly loweraflibercept recovery rate. The combination of DCM with 8% BSA presentedthe highest (97.26±5.37%) aflibercept recovery rate.

TABLE 3 Aflibercept Recovery Rate by BSA% Triacetin Ethyl Acetate DCM 0% BSA 90.97 ± 2.03 92.61 ± 4.61 87.85 ± 8.44  4% BSA 91.87 ± 6.2989.51 ± 4.99 90.95 ± 1.75  8% BSA 92.16 ± 6.35 91.45 ± 3.90 97.26 ± 5.3712% BSA 92.64 ± 2.93 88.69 ± 4.13 90.97 ± 3.22 16% BSA 86.26 ± 4.90 93.96 ± 12.57 86.80 ± 2.33 20% BSA 92.49 ± 3.93 81.61 ± 4.38  75.10 ±14.94

Optimal Mg(OH)₂ Loading for Stabilizing Aflibercept During Release forOne Month.

FIG. 17 summarizes the table below. As high as 19.64% of afliberceptstability was maintained at the end of first month for formulation 8%BSA+3% Mg(OH)₂ with DCM as organic phase.

TABLE 4 Aflibercept Stability Aflibercept Stability at end Formulationsof first month release 8% BSA + 3% Mg(OH)₂ in DCM 19.64% 12% BSA + 3%Mg(OH)₂ in DCM 2.74% (Original) 8% BSA + 0% Mg(OH)₂ in DCM 4.66% 8%BSA + 6% Mg(OH)₂ in DCM 1.95% 8% BSA + 9% Mg(OH)₂ in DCM 1.97%

Example 3

Even though anti-vascular endothelial growth factor (VEGF) therapy issuccessful for a majority of patients, there is a growing number ofpatients that do not respond to monthly monotherapy but do respond to acombination therapy such as corticosteroids and anti-VEGF. The examplebelow provides the extended and controlled dual release of dexamethasone(DEX) and aflibercept (AFL) from a single drug delivery system (DDS).

Two different preparations of single-emulsion poly(lactic-co-glycolicacid) (PLGA) nanoparticles were made by varying vortex and sonicationtime, resulting in DEX-A and DEX-B. Size distribution and mean diameterwere analyzed using Nanoparticle Tracking Analysis. The aflibercept wasencapsulated into PLGA microspheres using double emulsion technique. TheDDS for single release consisted of 20 mg of DEX nanoparticles (DEX-np)suspended within a biodegradable N-isopropylacrylamide/poly(ethyleneglycol)-co-(L-lactic acid) diacrylate/(NIPAAm/PEG-PLLA-DA)thermoresponsive hydrogel. The dual release consisted of 20 mg of DEX-npand 20 mg of AFL microspheres (AFL-ms) suspended within the hydrogel.DEX release in vitro, with and without AFL-ms, was quantified usingNanoDrop™ OneC. Iodine-125 radiolabeled AFL was used to measureencapsulation efficiency into the hydrogel and in vitro release. Theinitial burst was calculated by quantifying total drug release in thefirst 24 hours.

Average diameter was 138.9±6.2 and 267±55 nm for DEX-A and DEX-B,respectively. The single release of DEX-A and DEX-B had an initial burstof 288 and 301.2 ug, respectively. The addition of AFL-ms did notsignificantly alter the interval or steady state release ofdexamethasone for the first 60 days for DEX-A nor DEX-B. Conversely,increased release rates were seen for AFL in the presence of DEX-np inthe first 14 days. The addition of DEX-np reduced the encapsulationefficiency of AFL-ms into the hydrogel by 16.3%, the initial burst by1.3% and the final drug load dose by 7.9%.

DEX release kinetics from a hydrogel DDS were not significantly affectedby the presence of ALF-ms. AFL release rates from a hydrogel DDSincreased in the presence of DEX-np. This study suggests that anextended and controlled release of both DEX and AFL from a single DDScan be achieved.

Example 4

The varying loading doses of dexamethasone nanoparticles (20, 40, 60 and80 mg/ml) were also embedded into the hydrogel with afliberceptmicroparticles (20 mg/ml) to determine if the combination DDS wouldalter release characteristics. As seen in FIG. 18, the combination DDShad similar release rates for the first 140 days for all loading dosesof dexamethasone, resembling the single releases of dexamethasone.Around day 150, the 40 mg/ml loading dose had a steep increase inrelease rate. The increase in release rate was not seen in the higherdoses (60 and 80 mg/ml), as seen in Table 5.

TABLE 5 Release Characteristics of Dexamethasone Nanoparticles in DDSConcentration Estimated total Initial burst Release rate Release time ofDEX-np loading amount (first 24 after 7 days achieved (mg/ml) (μg)hours) (μg) (μg per day) (days) 20 3850.8 ± 99.8   244.51 ± 17.02 17.71224  40  6849.9 ± 1208.3* 294.91 ± 8.38 32.52 202* 60 5833.8 ± 134.9*323.83 ± 3.23 32.72 202* 80 8162.9 ± 288.2* 348.94 ± 6.53 41.08 202**Releases for 40, 60 and 80 mg/ml have estimated values for totalloading amount because the releases were ongoing.

FIG. 19 shows the cumulative releases of aflibercept microparticles (20mg/ml) with and without dexamethasone nanoparticles (20 mg/ml). Therelease of aflibercept from combination DDS had similar release kineticsas aflibercept from the single DDS but had a longer release time andhigher total amount released. The single release of aflibercept also hadsimilar release characteristics as previously reported. Table 6 showsthe release characteristics of aflibercept microparticles from singleand combination releases.

TABLE 6 Release Characteristics of Aflibercept Microparticles in DDSEstimated Initial E.E. of Release total loading burst (first particlesinto time amount 24 hours) hydrogel achieved (μg) (μg) (%) (days) SingleDDS 277.9 ± 8.4 66.7 ± 11.2 80.50 ± 5.8 203 Combination 267.4 ± 2.2 54.2± 7.7  74 67 ± 0.4 224 DDS

Example 5

Multi-drug release systems were prepared and tested according toembodiments of this invention. Sustained drug delivery systems (DDSs) ofthis invention can replace repeated injections and deliver awell-controlled drug release over a long period. The DDS systems of thisinvention protect drug bioactivity (such as protein-based drugs) forextended delivery application. Furthermore, the system can be modifiedto release multiple drugs.

A first example DDS could simultaneously release dexamethasone (DEX, acorticosteroid with 392 Da molecular weight) and aflibercept (AFL, 115kDa). DEX-nanoparticles (DEX-np) were fabricated using asingle-emulsion, solvent evaporation process. PLGA (75:25, MW 4-15 kDa)and DEX were dissolved in dichloromethane (DCM), with polyvinyl alcohol(PVA) used as the water phase. The mixture was sonicated at 100 wattsfor 3.5 minutes, and the solvent was evaporated to collect DEX-np, whichwas washed with deionized (DI) water at three times. The DEX-np werelyophilized to store. AFL-microparticle (AFL-mp) were fabricated usingthe double-emulsion, solvent evaporation method. A mix of DEX-np andAFL-mp was embedded into a NIPAAm/PEG-PLLA-DA thermoresponsive hydrogel.

The AFL-mp had an average diameter of 6.9±0.4 μm, and the encapsulationefficiency of the drug into microparticle was 66.6±1.4%. The DEX-np hadan average diameter of 138.9±6.2 nm and the encapsulation efficiency ofthe drug into nanoparticles was 91.6±4.2%. A mix of particle size ordrug characteristics (hydrophobic) did not alter the drug kinetics. Bothdrugs were able to release in a controlled manner for 224 days.Combo-DDS extended the release time of AFL by another 20 days (comparedto AFL-monotherapy DDS) and achieved a more complete release. Thebioactivity of drugs, and the treatment efficacy (using the CNV animalmodel), were both maintained through the 224 days, as shown in FIGS.20A-B.

A second example DDS could release growth differentiation factor 5(GDFS, aprotein with 27 kDa) and anti-tumor necrosis factor-alpha(anti-TNF alpha, Etanercept, ENT biologic with 150 kDa). DEX-np werefabricated using a single-emulsion, solvent evaporation process. PLGA(75:25, 1\4W 4-15 kDa) and DEX were dissolved in dichloromethane (DCM),with polyvinyl alcohol (PVA) used as the water phase. The mixture wassonicated at 100 watts for 3.5 minutes, and the solvent was evaporatedto collect DEX-np, which was washed with deionized (DI) water threetimes. The DEX-np were lyophilized to store. GDFS-microparticles(GDFS-mp) were fabricated using double-emulsion solvent evaporationusing 20 mg/ml stock. ENT-microparticles (ENT-mp) were fabricated usingdouble-emulsion, solvent evaporation method using 10 mg/ml stock. A mixof GDF5-mp and ENT-mp was embedded into the same hydrogel.

Two different biologics (large molecules) were encapsulated into twodifferent microparticles: GPF5-mp and ENT-mp, and the particles wereembedded into the hydrogel (Combo-DDS). FIGS. 21A-B show the releaseprofiles of single-DDS (either GDF5 or ENT) and Combo-DDS(GDF5-mp+ENT-mp-hydrogel). The presence of two different microparticlesdid not alter release kinetics. Releases were maintained for 175 days,and the drugs maintained bioactivity (in vitro model). Animal modelingof disc degeneration showed treatment efficacy (injection done betweenspinal discs).

A third example is a DDS where only hydrogel is used to releasevancomycin (VAN, antibiotics, 1450 Da). VAN was encapsulated in thehydrogel as a short-term release DDS (e.g., for use as a prophylactictreatment before ocular surgery). The release of VAN was controlled forover 20 days (see FIG. 22), and bioactivity was maintained throughoutthe release. Animal modeling (DDS injection was done via subconjunctivalinjection) showed treatment efficacy.

The size of nanoparticles and microparticles for this invention can becontrolled during the fabrication steps. The average nanoparticles usedwere approximately 140 nm, and average microparticles used wereapproximately 10 microns. Small molecules or hydrophobic drugs (likeDEX) are ideal for nanoparticles, as well as other agents likeantibiotics or other steroids. Microparticles are better suited in someembodiments for protein-based drugs and delicate biologics (AFL, GDF5,ENT), but other agents or cells can be encapsulated in microparticles.As discussed herein, the hydrogel itself is another source of drugrelease, such as for an extensive range of drugs (or even cells) and/orfor a shorter-term release (e.g., less than one month).

There are many combinations of microparticles, nanoparticles, andhydrogels that can be made according to this invention. As an example, amix of nanoparticles and microparticles in hydrogel are good forcombination of hydrophobic drugs and biologics. A mix of two differentmicroparticles in hydrogel is good for two different biologics. A mix oftwo different nanoparticles in hydrogel is good for two different smallmolecules.

Thus, the invention provides a new and improved method to delivertreatments for an extended period of time. Currently, there are over 4million intravitreal injections done monthly in US alone to treat AMD,DR and other vascular disease. Limiting injection to every 6 months willhave a great socioeconomic as well as reducing potential side effects ofmonthly intravitreal injections.

The invention illustratively disclosed herein suitably may be practicedin the absence of any element, part, step, component, or ingredientwhich is not specifically disclosed herein.

While in the foregoing detailed description this invention has beendescribed in relation to certain preferred embodiments thereof, and manydetails have been set forth for purposes of illustration, it will beapparent to those skilled in the art that the invention is susceptibleto additional embodiments and that certain of the details describedherein can be varied considerably without departing from the basicprinciples of the invention.

What is claimed is:
 1. A delivery composition, comprising a treatmentagent microencapsulated in degradable microcapsules that are suspendedin a degradable thermo-responsive hydrogel, wherein the degradablemicrocapsules comprise magnesium hydroxide (Mg(OH)₂) and bovine serumalbumin (BSA), wherein the hydrogel is thermo-responsive at aphysiological temperature of about 32° C. to about 37° C.
 2. Thecomposition of claim 1, wherein the microcapsules comprise: 0.001% to20% w/v BSA; and 0.001% to 9% w/v Mg(OH)₂.
 3. The composition of claim1, wherein the microcapsules further comprise polyethylene glycol (PEG)and sucrose.
 4. The composition of claim 3, wherein the microcapsulescomprise: 0.001% to 20% w/v BSA; 0.001% to 9% w/v Mg(OH)₂; 0.001% to 20%w/v PEG; and 0.001% to 10% w/v sucrose.
 5. The composition of claim 3,wherein the microcapsules comprise: about 10-14% w/v % BSA; about 2-4%w/v Mg(OH)₂; about 8-12% w/v PEG; and about 1.5-3.5% w/v sucrose.
 6. Thecomposition of claim 1, wherein the microcapsules further comprisepoly(lactic-co-glycolic acid), poly(lactic acid), polysaccharide chitin,alginate, or combinations or block copolymers thereof.
 7. Thecomposition of claim 6, wherein the microcapsules further comprisepolyethylene glycol (PEG) and sucrose.
 8. The composition of claim 1,further comprising a non-encapsulated treatment agent dispersed withinthe hydrogel.
 9. The composition of claim 1, further comprisingmicrocapsules having at least two release rates.
 10. The composition ofclaim 1, wherein the degradable microcapsules comprise a combination ofmicroparticles and nanoparticles.
 11. The composition of claim 10,wherein the microparticles contain a first treatment agent, and thenanoparticles contain a different second treatment agent.
 12. Thecomposition of claim 11, wherein each of the first and second treatmentagents is independently selected from an anti-VEGF agent, an anti-PDGFagent, cells, delivery cells, an antibiotic, a corticosteroid, enzymes,peptides, nucleic acids, or combinations thereof.
 13. The composition ofclaim 1, wherein the hydrogel comprises poly(N-isopropylacrylamide),poly(lactic acid), polysaccharide chitin, alginate, diacrylate, orcombinations or block copolymers thereof.
 14. A delivery composition,comprising: a degradable thermo-responsive hydrogel that isthermo-responsive at a physiological temperature of about 32° C. toabout 37° C. to provide a liquid-like state at room temperature and moresolid state at body temperature; degradable nanoparticles including afirst treatment agent, and suspended in the hydrogel; and degradablemicroparticles including a second treatment agent, and suspended in thehydrogel; wherein each of the degradable nanoparticles andmicroparticles comprises at least two of: polyethylene glycol (PEG),magnesium hydroxide (Mg(OH)₂), bovine serum albumin (BSA), and sucrose.15. The composition of claim 14, wherein the microcapsules furthercomprise: poly(lactic-co-glycolic acid), poly(lactic acid),polysaccharide chitin, alginate, or combinations or block copolymersthereof; 0.001% to 20% w/v BSA; 0.001% to 9% w/v Mg(OH)₂; 0.001% to 20%w/v PEG; and 0.001% to 10% w/v sucrose.
 16. A method of delivering acompound to an eye, the method comprising: applying to or into an eye ofa mammal a composition in a first physicochemical state, wherein thecomposition comprises a treatment agent microencapsulated in degradablemicrocapsules suspended in a degradable thermo-responsive hydrogel,wherein the degradable microcapsules are formed including at least twoof: polyethylene glycol (PEG), Mg(OH)₂, bovine serum albumin, andsucrose; and the composition changing to a second physicochemical stateupon administration, wherein the second physicochemical state is moresolid than the first physicochemical state, wherein the degradablemicrocapsules release the microencapsulated treatment agent over timeafter applying.
 17. The method of claim 16, wherein the microcapsulescomprise: 0.001% to 20% w/v BSA; and 0.001% to 9% w/v Mg(OH)₂.
 18. Themethod of claim 16, wherein the microcapsules comprise: 0.001% to 20%w/v BSA; 0.001% to 9% w/v Mg(OH)₂; 0.001% to 20% w/v PEG; and 0.001% to10% w/v sucrose.
 19. The method of claim 16, wherein the secondphysicochemical state is degradable to release the microencapsulatedtreatment agent and further comprising controlling the degradation by atleast one of: type and/or amount of crosslinking to control the releaseof the microencapsulated treatment agent, or selection of organicsolvent used in microencapsulation of the treatment agent.
 20. Themethod of claim 16, further comprising applying the composition in afirst physicochemical state by intravitreal injection, by periocular ortranscleral injection, by topical application, by intracameralapplication, by suprachoroidal application, within ocular implants, orcombinations thereof.